Rapid advances in biotechnology have led to the discovery of numerous protein and peptide therapeutics, many of which have recently reached the marketplace or are currently under regulatory review by the United States Food and Drug Administration. Unlike traditional small-molecule drugs, however, proteins and peptides generally cannot be administered orally; injection or infusion is most often required. Further, because of their fragility and short in vivo half-lives, encapsulation of proteins in biodegradable polymeric devices, from which the drug can be delivered, locally or systemically, for a prolonged period of time, has been a promising and intensely studied solution to these problems. Biodegradable microspheres comprising a variety of polymers have been the most studied devices due to relatively simple fabrication and facile administration to a variety of locations in vivo through a syringe needle.
Several methodologies for microsphere fabrication have been described, including precipitation, spraying, phase separation, and emulsion techniques. The emulsion and spraying approaches have been commonly used both at the bench and industrial scales. Sphere size and size distribution are reproducible but often poorly controllable. Standard deviations equal to 25-50% of the mean diameter are not uncommon.
Control of sphere size and size distribution has several important implications for controlled-release drug delivery. For example, there typically is an ideal sphere size that provides a desired release rate and route of administration. Spheres that are “too small” exhibit poor encapsulation efficiency, may migrate from the site of injection, and may exhibit undesirably rapid release of their payload. Spheres that are “too large” may not easily pass through a syringe needle. Thus, the typically polydisperse spheres generated by conventional fabrication techniques must be filtered or sieved to isolate particles within the desired size range, and the polymer and drug composing spheres outside that range are wasted.
Uniform microspheres approximately 1-5 μm in diameter would be ideal for passive targeting of professional antigen-presenting cells (APCs) such as macrophages and dendritic cells. Similarly, microspheres 10-20 μm in diameter could be used to target the tortuous capillary bed of tumor tissues by chemo-embolization. A system capable of precise microsphere fabrication could allow the optimal size for such applications to be identified and provide an efficient route to commercial manufacture and clinical implementation.
A long-sought goal for controlled-release drug delivery technologies is the ability to precisely control the release rate of encapsulated compounds, and microsphere size is a major determinant of release kinetics. Larger spheres generally release encapsulated compounds more slowly and over longer time periods, other properties (polymer molecular weight, initial porosity, drug distribution within the sphere, etc.) being equal. A constant (i.e., zero-order) release rate is often preferred, while variable drug release rates can be beneficial for many important indications. For example, intermittent high doses of antibiotics may alleviate evolution of resistance in bacteria, and discontinuous administration of vaccines often enhances the immune response.
Methods to control drug release rate include (i) choice of polymer chemistry (anhydrides, esters, etc.) and comonomer ratios, (ii) conjugating the drug to the polymer, (iii) varying the microsphere formulation parameters, and thus the physical characteristics of the resulting particles, and (iv) manipulating the sphere size and distribution. The success of the latter studies was limited by the relatively broad microsphere size distributions.
In recent years, there have been several reports of the fabrication of biodegradable polymer microspheres with controlled, uniform size (P. Sansdrap and A. J. Moes, Influence of manufacturing parameters on the size characteristics and the release profiles of nifedipine from poly(DL-lactide-co-glycolide) microspheres. Int. J. Pharm. 98 (1993) 157-164; B. G. Amsden and M. Goosen, An examination of the factors affecting the size, distribution, and release characteristics of polymer microbeads made using electrostatics. J. Control. Release 43 (1997) 183-196; K. Shiga, N. Muramatsu and T. Kondo, Preparation of poly(D,L-lactide) and copoly(lactide-glycolide) microspheres of uniform size. J. Pharm. Pharmacol. 48 (1996) 891-895; B. Amsden, The production of uniformly sized polymer microspheres. Pharm. Res. 16 (1999) 1140-1143; and N. Leelarasamee, S. A. Howard, C. J. Malanga and J. K. H. Ma, A method for the preparation of polylactic acid microcapsules of controlled particle size and drug loading. J. Microencapsul. 5 (1988) 147-157). However, none of these methods was successful in generating particles in a size range appropriate for drug delivery (˜1-100 μm) while maintaining narrow size distributions. In addition, these previous methods appear to be difficult to scale-up for commercial applications.
Hollow sphere fabrication techniques are disclosed in N. K. Kim, K. Kim, D. A. Payne, and R. S. Upadhye, “Fabrication of hollow silica aerogel spheres by a droplet generation method and sol-gel processing,” J. Vac. Sci., Technol. A., vol. 7, no. 3 pp. 1181-1184 (1989) and K. Kim, K. Y. Jang and R. S. Upadhye, “Hollow silica spheres of controlled size and porosity by sol-gel processing,” J. Am. Ceram. Soc., 74:8, pp.1987-1992, (1991).
Electrostatic spraying technique is disclosed in K. Kim and R. J. Turnbull, “Generation of charged drops of insulating liquids by electrostatic spraying,” J. Appl. Phys., vol. 47, no. 5, pp. 1964-1969, May 1976, U.S. Pat. No. 5,344,676 to Kim et al., and U.S. Pat. No. 6,060,128 to Kim, et al.
Previously developed techniques designed to fabricate hollow spheres employ a dual-nozzle scheme in which two coaxially mounted nozzles carrying different materials in liquid phase (the material in the inner nozzle could also be a gas) produce a smooth cylindrical jet which, in turn, is broken up into uniform droplets by an acoustic excitation. (See N. K. Kim, et al., “Fabrication of hollow silica aerogel spheres by a droplet generation method and sol-gel processing,” infra and K. Kim et al., “Hollow silica spheres of controlled size and porosity by sol-gel processing,” infra). The smallest drops that can be made with this method are roughly twice as large as the opening of the outer nozzle. This in turn indicates practical difficulties associated with fabricating uniform solid and hollow spheres of small sizes (less than about 50 μm in diameter) especially spheres in the submicron-size range. The reason is that the smaller the nozzle opening, the greater the chances for it to get plugged up, especially if the pharmaceutical compounds to be encapsulated are suspended as a particulate in the sphere-forming liquid. This problem becomes worse when the materials being used are viscous.
With previous technologies for spraying microdroplets from nozzle-type devices, the minimum sphere size typically obtainable is limited by the size of the nozzle opening. Usually, it is not possible to make drops smaller than the nozzle opening; typically, droplet diameters are 1-4 times the diameter of the nozzle. This presents several difficulties as the desired sphere size decreases. One problem is that fabrication of the nozzles themselves becomes more difficult as size decreases. This is especially true for large-scale fabrication methods in which it is necessary to form droplets through arrays of nozzles (perhaps 1000-2000). A second limitation stems from the pressure needed to pump fluids through small nozzles. The pressure required is given by
      Δ    ⁢                  ⁢    p    =            8      ⁢                          ⁢      μ      ⁢                          ⁢      L      ⁢                          ⁢      Q              π      ⁢                          ⁢              R        4            where Δp is the pressure drop across the nozzle, μ is the viscosity of the fluid, L is the length of the nozzle “passage”, Q is the volumetric flow rate of the fluid passing through the nozzle, and R is the radius of the nozzle opening. Thus, the pressure required scales with R−4. If one wishes to make microdroplets of ˜5 μm diameter, traditional methods may require a nozzle with a diameter of 5 μm or less. For example, at a flow rate of 1 mL/min and a fluid viscosity of 100 centipoise (100-times more viscous than water), a 5-μm diameter orifice would require a pump head of ˜1.1×1010 Pa (˜110,000 atm). This is clearly an impossibly high pressure. Even water, μ˜1 cp, requires a pressure of 1,100 atm to be pumped through a 5-μm diameter nozzle at 1 mL/min. Thus, pumping virtually any liquid through a nozzle of 5-μm diameter would require special equipment, if it could be done at all.
Another problem with traditional methods of forming small spheres is that some compounds to be encapsulated, such as plasmid DNA, may be damaged by shear forces. Damage depends on the product of the shear rate, γ, and the time spent in the shear field, θ. The average value of this product for a fluid flowing through a pipe is given by(γθ)avg=16/3·(L/D)where L is the length of the pipe and D is the pipe diameter. The orifice of a nozzle can be approximated as a pipe. However, entrance effects will tend to increase the shear rate meaning this equation will give a low estimate. Regardless, the value of γθ is approximately inversely proportional to the diameter of the orifice. Thus, decreasing the nozzle diameter from 100 to 5 μm would increase the damage done to any encapsulated compound by a factor of 20.
U.S. Pat. No. 6,116,516 describes stabilized capillary microjets, that produce aerosols. The microjets are formed by forcing a gas around a liquid stream. Under the correct conditions, micron-sized aerosols are produced, where preferably 90% or more have the same diameter plus/minus 3% to 30%.